1. Field of the Invention
The present invention generally relates to nuclear medicine, and systems for obtaining nuclear medicine images of a patient's body organs of interest. In particular, the present invention relates to systems that are capable of performing positron emission tomography (PET) as well as planar and single photon emission computed tomography (SPECT).
2. Description of the Background Art
Nuclear medicine is a unique medical specialty wherein radiation is used to acquire images which show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions which emanate from the body. One or more detectors are used to detect the emitted gamma photons, and the information collected from the detector(s) is processed to calculate the position of origin of the emitted photon from the source (i.e., the body organ or tissue under study). The accumulation of a large number of emitted gamma positions allows an image of the organ or tissue under study to be displayed.
Two basic types of imaging techniques are PET, and planar or SPECT imaging. PET imaging is fundamentally different from planar or SPECT imaging. In PET, events are detected from the decay or annihilation of a positron. When a positron is annihilated within a subject, two 511 KeV gamma rays are simultaneously produced which travel in approximately opposite directions. Two scintillation detectors are positioned on opposite sides of the patient such that each detector will produce an electrical pulse in response to the interaction of the gamma rays with a scintillation crystal. In order to distinguish the detected positron annihilation events from background radiation or random events, the events must be coincident in each detector in order to be counted as "true" events. When a true event is detected, the line connecting the positions of the two points of detection is assumed to pass through the point of annihilation of the positron.
By contrast, single photon imaging, either planar or SPECT, relies on the use of a collimator placed in front of a scintillation crystal or solid state detector, to allow only gamma rays aligned with the holes of the collimator to pass through to the detector, thus inferring the line on which the gamma emission is assumed to have occurred. Both PET and single photon imaging techniques require gamma ray detectors that calculate and store both the position of the detected gamma ray and its energy.
PET imaging systems and single photon imaging systems have been known and commercially available for many years. Recently, manufacturers of single photon imaging systems have modified such systems to enable them to perform PET imaging by adding the capability to detect coincident events (i.e., two gamma ray interactions in opposing detectors occurring within a small time interval).
Present day single photon imaging systems all use large area scintillation detectors made of sodium iodide crystals doped with thallium, or NaI(Tl). NaI(Tl) is well suited for the detection of lower energy (e.g., 140 keV for .sup.99m Tc and about 72 keV for x-rays from .sup.201 Tl) single photon emitting radioisotopes, but it is not the optimum material for detecting high energy (e.g., 511 keV) positron annihilation isotopes.
In order to efficiently detect 511 keV photons from a positron annihilation, the detector material should have a high stopping power to maximize the probability of interaction of the photon in the detector, and high effective Z (atomic number) to maximize the probability that the gamma photons will be absorbed via the photoelectric effect rather than being scattered via the Compton scattering effect.
Another performance problem with NaI(Tl) detectors is the relatively low count rate capability of such systems. The count rate is the ability to detect and resolve independent gamma events occurring within one second. NaI(Tl) detectors current cannot resolve independent gamma events in excess of a few hundred thousand per second. In the case of PET imaging using NaI(Tl) detectors, the number of gamma pairs detected that are in true coincidence can be as low as 1%. Therefore the true coincidence count rate of a NaI PET detector system is severely limited (.about.10 k counts/sec) as compared with a dedicated PET imaging system.
Another known hybrid PET/SPECT system uses two relatively new scintillating materials, namely LSO (lutetium oxyorthosilicate) and YSO (yttrium oxyorthosilicate). In this system, the relatively high Z LSO crystal is placed behind the lower Z YSO crystal, forming a sandwich of scintillators, sometimes called a phoswich detector. When carrying out PET imaging, many high energy gammas passing through the low Z detector will be absorbed in the high Z LSO detector, while when performing single photon imaging, a collimator is placed in front of the crystal, and most of the low energy gamma photons will be absorbed in the low Z YSO crystal.
The count rate capability of this proposed system is much higher than NaI because of the faster decay time of the LSO and YSO detector materials as compared with NaI(Tl), and the PET performance is better than NaI because of the higher stopping power and higher effective Z of the LSO material. The single photon imaging capability is provided by the YSO layer, which is slightly inferior to NaI in effective Z, but slightly superior in total light output.
Another recent development in the art is the use of CZT (cadmium zinc telluride) as a solid state (i.e., semiconductor) detector material. As a single photon detector, CZT is superior to NaI in several performance parameters. First, its energy resolution is less than 5%, as compared with 9-10% for NaI. The effective Z of CZT is about the same as NaI, but its density is higher, making its stopping power about 30% better than NaI per unit thickness of material. Third, the count rate capability for CZT detectors is virtually unlimited as compared to a scintillator crystal, because each pixel (or picture element) of the CZT material acts as an independent detector. Thus, unlike a scintillator crystal, in which two events occurring very close in time and spatial location will produce overlapping light output, two gamma photons arriving at exactly the same time in adjacent pixels of a CZT detector could be independently detected and measured accurately with respect to energy, given an optimum electronic circuitry design.
With respect to PET imaging, CZT is not more attractive than NaI in terms of effective Z, but would have a higher count rate capability. In order to maximize the probability of detection, the thickness of the CZT detector should be as large as possible. Thicker material, however, would increase the charge collection time and degrade the temporal resolution which is important for coincidence detection. Additionally, the high cost of CZT material as compared with NaI may prove prohibitive.
There thus remains a need in the art for a system capable of performing both PET imaging and single photon (planar or SPECT) imaging, separately or simultaneously, without comprising accepted image quality requirements.